JP5159145B2 - Shield coil and magnetic resonance imaging apparatus - Google Patents

Description

The present invention relates to a shield coil and a magnetic resonance imaging apparatus for shielding a magnetic field generated in a magnetic resonance imaging apparatus , and more particularly to a passive shield coil and a magnetic resonance imaging apparatus formed of a superconducting material.

  A medical magnetic resonance imaging (MRI) apparatus includes a static magnetic field coil that generates a static magnetic field, a shim coil that corrects the uniformity of the static magnetic field, a gradient magnetic field coil that generates a gradient magnetic field superimposed on the static magnetic field, and transmission and reception of high-frequency signals. Various magnetic field generating coils such as RF coils are used.

  Among these, the shim coil, the gradient magnetic field coil, and the RF coil are arranged inside a magnetic field space (for example, a cylindrical space) as a diagnostic space formed by a static magnetic field coil (for example, a superconducting magnet). In this case, normally, the shim coil is disposed at a position closest to the static magnetic field magnet, and then the gradient magnetic field coil is disposed. Conversely, the RF coil is disposed at a position closest to the subject inserted in the diagnostic space.

  As the shim coil, a normal conducting shim, a superconducting shim, and a passive shim are frequently used. In addition, in gradient magnetic field coils, the shield type is used to prevent the gradient image pulse generated according to the pulse sequence from generating eddy currents in the housing that supports the static magnetic field coil and the like, thereby reducing the image quality. The coil is used frequently. From this point of view, the gradient magnetic field coil includes an active shield type gradient magnetic field coil and a passive shield type gradient magnetic field coil.

  Of these, the active shield type gradient magnetic field coil generally includes a main coil that generates a gradient magnetic field pulse having a desired spatial distribution, and a shield coil disposed around the main coil. Leakage to the outside of the gradient magnetic field pulse generated by the main coil is suppressed by the shielding magnetic field that is actively generated by the shield coil. For this reason, the gradient magnetic field leaking to the outside of the coil unit is reduced, and the influence of eddy currents on imaging due to the leaked magnetic field is suppressed.

  On the other hand, the passive shield type gradient magnetic field coil includes a main coil that generates a gradient magnetic field pulse having a desired spatial distribution, and a shield body disposed around the main coil. This shield body is made of, for example, a cylindrical member made of a shield material, and is shielded so as to reduce leakage of gradient magnetic field pulses to the outside without being connected to a current source unlike the active type.

  In general, the active shield type gradient magnetic field coil is superior in terms of shielding performance, but the structure is more complex than the passive shield type, and the design and manufacturing accuracy of the shield coil pattern is improved. Higher ones are required.

By the way, what was described in patent document 1 is known as a shield between a gradient magnetic field coil and a static magnetic field coil in recent years. According to the shield structure described in this document 1, as shown in FIG. 1, in a magnetic resonance imaging apparatus using a superconducting magnet device, both a gradient magnetic field coil and a current supply lead to this coil are superconducting. Each is covered with a magnetic shield made from a multilayer composite. As a result, it is possible to reduce the striking sound caused by the electromagnetic force caused by the pulse current flowing through the gradient magnetic field coil and its current supply lead placed in the static magnetic field. Further, according to this configuration, it is considered possible to reduce the leakage of the gradient magnetic field pulse generated by the gradient magnetic field coil to the outside and reduce the eddy current due to the leakage magnetic field.
JP-A-8-84712

  However, according to the shield structure described in Patent Document 1, for example, a superconducting multilayer composite having a thickness of about 1 mm in which an NbTi layer (30 layers), an Nb layer (60 layers), and a Cu layer (31 layers) are multilayered. Since it is formed, it has a shape of a superconducting sheet in which a superconducting substance is distributed over the entire surface. For this reason, depending on the distribution of the applied static magnetic field, the superconducting sheet is exposed to a very strong magnetic field, and the shield current flowing in the superconducting material may be saturated by the Meissner effect. . For this reason, when this saturation occurs, the critical point is reached in many parts of the superconducting sheet and becomes a normal state, and the heat generation at this time causes the entire sheet to be in a normal state and the desired shielding effect cannot be obtained. There is. In addition, the production of this superconducting sheet involves processes such as hot rolling and cold rolling by laminating multiple members, which makes the production cost very high. It is difficult in terms of cost.

Therefore, the present invention was made in view of the shield structure formed of the above-described superconducting multilayer composite, and even when used in a magnetic resonance imaging apparatus, it prevents a decrease in shield performance due to the Meissner effect, Passive suitable for magnetic resonance imaging that provides reliable and stable shielding performance against the desired shield target, while being relatively easy and can be manufactured at a lower cost by using less superconducting material It is an object of the present invention to provide a shield coil of a type and a magnetic resonance imaging apparatus .

In order to achieve the above-described object, according to one aspect of the present invention, there is provided a passive shield coil that shields a magnetic field generated from a magnetic field generating coil provided in a magnetic resonance imaging apparatus. The body is formed by a portion formed of a superconducting material and a portion without the superconducting material, and the portion without the superconducting material is a portion due to lattice defects of the superconducting material.

According to the present invention, even when a shield coil formed of a superconducting material is used in a magnetic resonance imaging apparatus, it is possible to prevent a decrease in shield performance due to the Meissner effect and to reliably and stably protect a desired shield target. On the other hand, it is possible to provide a passive type shield coil and a magnetic resonance imaging apparatus which can be manufactured more inexpensively while exhibiting the above shielding performance, and relatively easily and using less superconducting materials.

  An embodiment of a shield coil according to the present invention will be described.

  With reference to FIGS. 1-8, one embodiment of the shield coil based on this invention is described. This shield coil is a passive shield coil that does not require power supply, and is integrated with the gradient coil to form a passive shield gradient coil unit. Since this gradient magnetic field coil unit is provided in the magnetic resonance imaging apparatus, the magnetic resonance imaging apparatus will be described.

  FIG. 1 shows a schematic cross section of a gantry 1 of a magnetic resonance imaging (MRI) apparatus. The entire gantry 1 is formed in a cylindrical shape, and the bore at the center functions as a diagnostic space. At the time of diagnosis, the subject P can be inserted into the bore. The gantry 1 is connected to an imaging apparatus that performs processing such as application of a gradient magnetic field and a high-frequency signal, collection of echo signals, and image reconstruction along a pulse sequence, thereby enabling magnetic resonance imaging. Yes.

  The gantry 1 includes a substantially cylindrical static magnetic field coil unit 11, a substantially cylindrical gradient magnetic field coil unit 12 disposed in the bore of the coil unit 11, a shim coil unit 13 attached to, for example, the outer peripheral surface of the unit 12, And an RF coil 14 disposed in the bore of the gradient coil unit 12. The subject P is placed on a couch top (not shown) and loosely inserted into a bore (diagnostic space) formed by the RF coil 14.

The static magnetic field coil unit 11 is formed of a superconducting magnet. That is, a plurality of heat radiation shield containers and a single liquid helium container are housed in an outer vacuum container, and a superconducting coil is wound and installed inside the liquid helium container.

  Here, the gradient magnetic field coil unit 12 is formed in a passive shield type as described above. The coil unit 12 has a coil assembly for each of the X, Y, and Z channels in order to generate a pulsed gradient magnetic field for each of the X-axis direction, the Y-axis direction, and the Z-axis direction. Moreover, the entire unit has a passive shield structure that hardly leaks the gradient magnetic field of each channel to the outside in the unit radial direction.

  Specifically, as shown in FIG. 2, in this gradient coil unit 12, cylindrical main coils (gradient coils) 12X, 12Y, 12Z of X, Y, Z channels are provided for each coil layer. Laminated while being insulated. Further, one cylindrical shield coil that collectively shields the pulsed gradient magnetic fields generated by the main coils 12X, 12Y, and 12Z of all the channels, that is, the gradient magnetic fields that change dynamically, outside the outermost main coil 12Z. 12S is arranged.

  Of these, as shown in FIG. 3, the Y-channel main coil 12Y includes four saddle-shaped spiral winding portions CR1, CR2, CR3, and CR4 each having a flat conductor wound around a small-diameter bobbin B1. Is provided. Winding portions CR1 and CR2 are juxtaposed in the Z-axis direction and electrically connected in series. A pair of winding portions composed of the winding portions CR3 and CR4 are arranged to face each other at a position where the pair of winding portions is rotated 180 degrees with respect to the Z axis. The direction of the pulse current flowing through each winding part is set as shown by the arrows in the figure between the opposing winding parts and the winding parts arranged in parallel so that the magnitude of the gradient magnetic field in the Y-axis direction changes linearly. ing.

  FIG. 3 schematically shows a saddle-like spiral winding pattern for one turn (winding), and each actual winding portion is composed of a plurality of turns of a saddle-like spiral coil. The bobbin B1 is formed by winding an insulating layer on a winding layer of another channel.

  The X-channel main coil 12X is similarly arranged in a state where the Y-channel main coil 12Y is rotated 90 degrees with respect to the Z-axis.

  Further, as schematically shown in FIG. 4, the Z-channel main coil 12Z is formed by winding a flat conductor around a cylindrical bobbin B2 in a plurality of turns along a solenoidal coil pattern. In the main coil 12Z, the winding direction is opposite to the center position Z = 0 in the Z-axis direction.

The shield coil 12S has a characteristic structure implementing the present invention, and forms a passive shield coil made of a superconducting material. Specifically, as shown in FIG. 5, the shield coil 12S is composed of a mesh cylindrical body (coil body) 21N formed by using, for example, an NbTi-based superconducting wire 21 as a superconducting material. That is, the cylindrical body (coil body) 21N has a simple configuration in which the superconducting wire 21 is assembled so as to draw a rhombus pattern along the peripheral surface. For this reason, the shield coil 12S is configured by a plurality of substantially rhombus frame portions 21A made of a superconductor formed along the peripheral surface as the coil portion. The inside of each of the plurality of substantially diamond-shaped frame portions 21A is a hole (space portion) 21B.

  That is, the “hole” refers to a portion (region) where no superconducting material exists. For this reason, any portion without superconducting material may be used, so it may or may not be an actual “hole”. For example, the “hole” may be filled with a material having good thermal conductivity and conductivity such as copper, for example, a normal conductive material. Further, such a “hole” may be a portion having no superconducting material, and may be, for example, a very small spatial hole at an atomic level called a lattice defect. A superconducting material having such lattice defects, for example, a superconducting sheet, can be produced with another purpose by adding impurities to increase the current density. In addition, in order to form the shield coil according to the present invention as a dynamic shielding sheet that does not generate heat (not necessarily a sheet), at least the spatial superconducting hole is surrounded by a superconductor. It is desirable. This is because it is desired to pass a shielding permanent current to shield the dynamic magnetic field fluctuation.

  As described above, the shield coil according to the present invention having a portion without the superconducting material exhibits a completely different behavior from the conventional shield coil using the superconducting sheet filled with the superconducting material. The behavior will be described later with reference to specific drawings, but here is simply described as follows. In the shield coil according to the present invention, by providing a place where the superconducting material does not exist in the space, regarding the static magnetic field, it behaves as if there is no superconducting material, and regarding the dynamic magnetic field change, A shielding permanent current is generated on the superconducting material that cancels the magnetic flux passing through the hole without the superconducting material, selectively shielding only the dynamic magnetic field change. At this time, an eddy current is generated on a conductive conductor such as copper which is physically filled for the purpose of protecting the superconducting material, but the current is immediately absorbed by the superconducting material having zero electrical resistance. , The heat generated by it is very small.

  Returning to the description of the shield coil 12S according to the embodiment again. The shield coil 12S is housed in a cylindrical container 22 and can be filled with liquid helium to bring the superconducting wire 21 into a superconducting state.

  The shield coil 12S can be manufactured by various methods. 6 and 7 show an example of such a manufacturing method. In any of these manufacturing methods, a superconducting wire 21 in which a superconducting material (core material) made of, for example, an NbTi-based superconducting material is embedded in the conductor along the length direction thereof is used.

  In the case of the manufacturing method shown in FIG. 6, a plurality of such superconducting wires 21 are arranged in parallel and electrically connected at predetermined intervals between adjacent wires (formation of electrical junctions CN) (FIG. 6). (See (A)). Next, the wire formed with a plurality of electrical junctions CN in this manner is stretched so as to draw a diamond-shaped pattern (hole 21B) having four electrical junctions CN as vertices (see FIG. B)). Next, the stretched wire is rounded into a cylindrical shape along the cylindrical bobbin B3 (see FIG. 5).

  As another example of this manufacturing method, cuts of a predetermined length are put in a plurality of rows at predetermined intervals in a superconducting sheet whose entire surface is formed of a superconducting material, and then the sheet is spread laterally. You may do it. As a result, each of the cut portions expands to form a hole portion 21B, and each of the connection portions of the cut portions forms the electrical junction CN described above, and the shield coil similar to that of FIG. 12S can be formed.

  On the other hand, in the manufacturing method shown in FIG. 7, the superconducting wire 21 is first helically wound at a predetermined oblique angle along the peripheral surface of the prepared cylindrical bobbin B3. Next, the direction is changed so as to intersect the winding direction of the superconducting wire, and the superconducting wire 21 is wound helically again. Due to the two helical windings, the positions where the superconducting wires 21 intersect along the peripheral surface form the electrical junction CN described above, and the substantially rhombus pattern portion surrounded by the electrical junction CN. Forms the hole 21B.

  Next, the effect of the shield coil 12S of the gantry 1 of the magnetic resonance imaging apparatus of the present embodiment will be described.

  The purpose of the shield coil 12S is to prevent leakage of the magnetic field from the gradient magnetic field coil unit 12 to the direction of the static magnetic field coil unit 11 (outward direction of the radius of the gantry 1). For this reason, it is necessary to shield a gradient magnetic field (dynamically changing magnetic field) that changes in a pulse shape without exhibiting a shielding effect against a static magnetic field. For this purpose, it is important to manage the order of application of the static magnetic field when the shield coil 12 (ie, the superconducting wire 21) is transitioned to the superconducting state. That is, it is important to apply a static magnetic field first and then transition the shield coil 12 to the superconducting state.

Therefore, when magnetic resonance imaging is executed using this magnetic resonance imaging apparatus, the static magnetic field coil unit 11 is first activated to generate a static magnetic field in accordance with the order relation. Since this static magnetic field generates a static magnetic field that passes through the inside of the gantry 1, the magnetic flux B ₀ of this static magnetic field passes through the shield coil 12 as shown in FIG. That is, since the shield coil 12 at this time is not yet in the superconducting state, the magnetic flux B _{0 of the} static magnetic field can freely pass through the hole 21B.

  Next, after this static magnetic field is generated, the shield coil 12 is transitioned to the superconducting state. For this purpose, the container 22 may be filled with liquid helium. By this filling, the shield coil 12 is immersed in liquid helium in the container 22. As a result, the shield coil 12 is cooled to an extremely low temperature of liquid helium: −269 ° C., and thus transitions to a superconducting state.

  Thereafter, the X, Y, and Z channel gradient magnetic field coils (main coils) 12X, 12Y, and 12Z of the gradient magnetic field coil unit 12 and the RF coil 14 are driven and echoed along a predetermined pulse sequence. A signal is collected via the RF coil 14.

The gradient magnetic field coils (main coils) 12X, 12Y, and 12Z are driven by a pulse signal that is turned on / off within a very short time. Therefore, the gradient coils 12X, 12Y, and the gradient G generated from 12Z (Gx, Gy, Gz) gives the spatial position information to the static magnetic field _{B 0} being superimposed on the static magnetic field _{B 0.} Therefore, it is possible to collect the desired echo signal to the original static field B _0.

Each gradient coil 12X, gradient G generated 12Y, and the 12Z, while superimposed on the static magnetic field B _0, is radiated also to the side of the static magnetic field coil unit 11. This gradient magnetic field G is a magnetic field that dynamically changes with time.

For this reason, the magnetic flux of the gradient magnetic field G tends to pass through the hole 21B because the frame portion 21A of the shield coil 12S is completely non-magnetic. However, as shown in FIG. 8, since each of the holes 21B is a single closed loop, the loop current induced in each closed loop cancels out between adjacent loops. Not flowing. That is, although the static magnetic field B ₀ which has been initially applied to pass through the shield coil 12S, the magnetic field applied after the transition to the subsequent superconducting state, that is, it can not pass through the magnetic field gradient G shield coil 12S, It is shielded by the shield coil 12S. Therefore, since it does not participate in the static magnetic field B ₀ at all, a high-performance shielding effect of selectively shielding only the gradient magnetic field G without reaching a critical point by the static magnetic field and causing a quench phenomenon can be obtained.

  For this reason, the gradient magnetic field G leaks to the static magnetic field coil unit 11 side, an eddy current is generated in the casing of the unit 11, and the image quality of the reconstructed MR image is deteriorated by this eddy current. It can be reliably eliminated or reduced.

  In addition, in the case of the passive shield coil 12S according to the present embodiment, the dynamic variation magnetic field (gradient magnetic field) of all the X channel, the Y channel, and the Z channel is shielded while being one (one layer) shield. be able to. That is, the shield can be shielded by one passive shield regardless of the shape and winding pattern of the gradient magnetic field coils 12X, 12Y, and 12Z located on the inner diameter side (bore side) of the shield coil 12S. For this reason, the shield coil 12S can be used as a universal shield excellent in versatility.

  Further, as described above, since it can be applied to gradient magnetic field coils for various uses and it is not necessary to supply current, the shield coil is contained in the superconducting magnet container (vacuum container) employed with the static magnetic field coil unit 11. 12S can also be arranged. By arranging in this way, the distance between the gradient coils 12X, 12Y, and 12Z, which are the main coils, can be increased. Therefore, the capacity of the power source supplied to the main coil can be reduced, the efficiency of the gradient coil unit 12 can be further improved, and further downsizing can be achieved.

  Furthermore, since the superconducting material used for the shield coil 12S only needs to be a wire such as NbTi wire, unlike the conventional superconducting sheet type shield coil, the cost of the superconducting material can be reduced and manufactured at a lower cost. In practice, a very important effect can be obtained.

  Further, a mechanism (the container 22 and a port for the same) for transitioning the passive shield coil 12S to the superconducting state is necessary, but unlike the active type gradient magnetic field coil unit, one layer common to each channel. In consideration of the fact that the shield layer (shield coil 12S) is sufficient and that no power supply or wiring is required for the shield coil 12S, the passive type gradient magnetic field coil unit 12 using the shield coil 12S The overall configuration is simplified and miniaturized.

  By the way, since the shield coil 12S according to the present embodiment is allowed to transit to the superconducting state after generating the static magnetic field, the static magnetic field can be continuously passed. On the other hand, in the case of a superconducting sheet in which a superconducting material is simply distributed over the entire surface as in the past, due to the Meissner effect, the magnetic flux is driven out of the sheet regardless of the order in which the static magnetic field is generated and applied. Therefore, the present embodiment is decisively different from the present embodiment in that both the static magnetic field and the variable magnetic field are shielded.

  As described above, when the container 22 for filling liquid helium is not provided, a direct-attached type refrigerator is used, and the shield coil 12S is shifted to the superconducting state by starting the refrigerator. May be. For this reason, various aspects can be taken also about the structure for control from a normal conduction state to a superconducting state.

(Another embodiment)
Furthermore, another embodiment according to the shield coil of the present invention will be described.

  The shield coil according to this embodiment is implemented for the shim coil unit 13 of the magnetic resonance imaging apparatus.

  As shown in FIG. 9, the shim coil unit 13 includes a shield coil 13 </ b> S having a mesh structure similar to the shield coil 12 </ b> S illustrated in FIG. 5 described above. The shield coil 13S further includes a current supply circuit 23 that can be turned on / off, and a switch 24 that turns the current supply circuit 23 on / off. Therefore, when the switch 24 is operated to turn on the current supply circuit 23, a current is supplied to the shield coil 13S, and the coil 13S is heated.

  When the shield coil according to the present invention is implemented as a shimming coil as in the present embodiment, it is necessary to shield not only the dynamically changing gradient magnetic field G but also the static magnetic field as described above. is there. For this purpose, the timing of application of a static magnetic field and the timing of transition of the shield coil 13S to the superconducting state becomes important as in the above-described embodiment. In the case of the above-described embodiment, it is necessary to shift the shield coil 13S to the superconducting state after applying the static magnetic field, but in this embodiment, the order is reversed.

  A specific example of this is as follows. Now, it is assumed that the shield coil 13S is already in a superconducting state by filling the container 22 with liquid helium (at this time, the static magnetic field coil unit 11 has not been started yet). In this state, the switch 24 is operated to supply current from the current supply circuit 23 to the shield coil 13S. Thereby, since the shield coil 13S is heated by Joule heat, the shield coil 13S returns to the normal conduction state instantaneously.

  As described above, once the state is surely returned to the normal conduction state, the switch 24 is operated to stop the current supply from the current circuit 23 to the shield coil 13S. For this reason, the shield coil 13S immersed in the liquid helium in the container 22 again transitions to the superconducting state.

Next, the static magnetic field coil unit 11 is already activated to generate a static magnetic field. Accordingly, the magnetic flux of the static magnetic field B ₀ is already can not pass through the shield coil 13S which is a superconducting state, also, the magnetic flux of the gradient G is also not allowed to pass through the shield coil 13S. That is, both are shielded by the shield coil 13S.

By disposing such a shield coil 13S, in addition to the blocking function of the gradient G, shimming effect due to the static magnetic field B ₀ in the shielding function can be obtained.

  The gradient magnetic field coil unit 12 (shield coil 12S) and the shim coil unit 13 (shield coil 13S) may be used in combination.

  In the shield coil 12S described above, the superconducting wire 21 is used to form a plurality of substantially rhombic holes 21B. Basically, however, the shape of the holes 21B is not necessarily limited. It is not limited to this.

  In other words, the biggest concept of this hole is that the superconducting material does not have to be distributed over the entire surface of the shield coil, and a plurality of regular and / or irregular holes are formed. As long as the magnetic flux of the static magnetic field can pass therethrough. As a mode of forming the regular hole portion, there is formation of a mesh-shaped hole portion having a fixed shape.

  The hole 21B having a fixed shape is not necessarily limited to a rhombus, and as shown in FIGS. 10A to 10C, a circle, an oval, a square, an ellipse, an oval, a rhombus, etc. Any shape is acceptable. Each of the holes 21B functions as a small closed loop coil, and can shield the magnetic flux entering the coil. For example, in the case of the circular hole 21B as shown in FIG. 10A, a plurality of small loop coils are manufactured using a wire material such as NbTi wire, and these loop coils are connected on the circumferential surface of the bobbin. By arranging, the shield coil 12S can be formed.

  Moreover, even if it is the hole part 21B of the same fixed shape, you may make it adjust the magnitude | size according to the position of the cylindrical shield coil 12S (13S). For example, as schematically shown in FIG. 5, the size of the hole 21B is made fine at the center of the shield coil 12S in the direction of the cylindrical axis (Z-axis), and on the contrary, it is closer to the end on the cylindrical axis. The hole 21 </ b> B may be formed rougher as it becomes. Thereby, the shield performance can be adjusted according to the importance of the position, for example, the shield performance can be improved as it approaches the central portion.

  Furthermore, even if the superconducting material is a shield coil having a structure in which a plurality of small coils are formed via a normal conducting material, the same shielding effect is exhibited. In this case, however, there is a problem of a quenching phenomenon due to heat generation, so it is desirable that the superconducting material is distributed in a linear form to form a coil.

  Furthermore, you may form the shield coil which concerns on the Example and modification which were mentioned above as a multilayer structure of at least 1 layer or more. Moreover, you may arrange | position the shield coil which concerns on the Example and modification which were mentioned above in the same vacuum layer as another static magnetic field coil which generate | occur | produces a static magnetic field, or another independent vacuum layer.

  Furthermore, the above-described gradient coil may be formed as a so-called self-shielding gradient coil (for example, Actively Shielded Gradient Coil (ASGC)), and the shield coil described above may be applied thereto. . Further, the shield coil according to the present invention can be implemented as a non-shield type gradient magnetic field coil. Furthermore, in order to partially shield the gradient coil, one or more shield coils according to the present invention may be arranged outside the gradient coil.

  The present invention is not limited to the above-described embodiments and modifications thereof, and those skilled in the art will be able to use various conventionally known techniques within the scope of the gist of the claims. Modifications can be made and these are within the scope of the present invention.

1 is a schematic cross-sectional view around a gantry of a magnetic resonance imaging apparatus according to an embodiment of the present invention. FIG. 2 is a schematic sectional view taken along line II-II in FIG. 1. The schematic block diagram of the gradient magnetic field coil of a Y channel. The schematic block diagram of the gradient magnetic field coil of a Z channel. The schematic block diagram including the partial expansion explanatory drawing of the shield coil common to X, Y, and Z channel. The figure explaining an example of the manufacturing method of a shield coil. The figure explaining the other example of the manufacturing method of a shield coil. The figure explaining the shield operation of the shield coil in an embodiment. The schematic block diagram explaining the shield coil which concerns on other embodiment of this invention. The figure explaining the modification of a shield coil and its hole part.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 Gantry 11 of magnetic resonance imaging apparatus Static magnetic field coil unit 12 Gradient magnetic field coil unit 13 Shim coil unit 14 RF coil 12X, 12Y, 12Z Gradient magnetic field coil (main coil)
12S shield coil 21 superconducting wire 22 container 21A frame 21B hole 21N cylindrical body (coil body)
23 Current supply circuit 24 Switch

Claims (10)

A passive shield coil that shields a magnetic field generated from a magnetic field generating coil provided in a magnetic resonance imaging apparatus,
A shield coil , wherein the coil body is formed of a portion formed of a superconducting material and a portion without the superconducting material, and the portion without the superconducting material is a portion due to lattice defects of the superconducting material. . The magnetic field generating coil is a three-channel gradient magnetic field coil including an X channel, a Y channel, and a Z channel of the magnetic resonance imaging apparatus, and is disposed at a predetermined position on the outer peripheral side of the gradient magnetic field coil. The shield coil according to claim 1, wherein the gradient magnetic field of each channel generated by the gradient magnetic field coil is shielded from leaking to the outer peripheral side. The magnetic field generating coil is a static magnetic field generating means that is provided in the magnetic resonance imaging apparatus and generates a static magnetic field and has a bore on the inner peripheral side, and a predetermined position on the inner peripheral side of the static magnetic field generating means The shield coil according to claim 1, wherein the shield coil is arranged to be responsible for homogenization of the static magnetic field. The shield coil according to any one of claims 1 to 3 , wherein the coil body is formed in a multilayer structure of at least one layer or more. The shield coil according to any one of claims 1 to 4 , wherein the coil body is arranged in the same vacuum layer as a static magnetic field coil that generates a static magnetic field. The shield coil according to any one of claims 1 to 4 , wherein the coil body is disposed in an independent vacuum layer different from a static magnetic field coil that generates a static magnetic field. The shield coil according to claim 2 , wherein the gradient coil is formed as a self-shielding gradient coil and applied to the gradient coil. The shield coil according to claim 2 , wherein the gradient coil is formed as an unshielded gradient coil and applied to the gradient coil. The shield coil according to claim 2 , comprising at least one coil body disposed outside the gradient magnetic field coil and disposed to partially shield the gradient magnetic field coil. A magnetic resonance imaging apparatus comprising the shield coil according to claim 1 .

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